There exist several medical therapies that involve inducing temperature changes in parts of the body. To facilitate the monitoring of such therapies, it would be desirable to provide apparatus and techniques for non-invasively measuring the temperature of an internal treatment site.
One such therapeutic technique involves the use of high intensity ultrasound fields, which can be focused on deep seated regions within the human body. If sufficient acoustic energy is concentrated within the focal volume, the temperature in that region can be increased to a level sufficient to induce cellular necrosis. This technique is commonly referred to as High Intensity Focused Ultrasound (HIFU) or Focused Ultrasound Surgery (FUS). The focusing of ultrasound can result in high acoustic intensities (measured as power density in W/cm2) at the focus. An intensity gain of 100 to 1000 times can be achieved in the cross-sectional area of the beam at the focus, resulting in intensities upward of 1000 W/cm2, orders of magnitude greater than that of diagnostic ultrasound systems. The technique has potential applications in any medical field that may benefit from the selective destruction of tissue volumes. Clinical interest has concentrated on treating soft-tissue cancers, de-bulking enlarged prostates (benign prostatic hyperplasia), and acoustic homeostasis. HIFU provides the ability to ablate localized tissue volumes without damaging intervening and surrounding tissue, thus eliminating the need for incisions—a feature that distinguishes it from other widely used ablative therapies, such as Radio Frequency (RF) ablation.
HIFU therapy is significantly different than conventional hyperthermia treatment, in which the temperature is raised a few degrees above body temperature and maintained for a relatively lengthy time, typically on the order of minutes. In HIFU therapy, the temperature typically rises by more than 40-50° C. in a few seconds, with increases up to 70° C. in one second having been reported. One consequence of the rapid heating to cytotoxic temperatures is that there is a very narrow boundary between live cells and dead cells at the edges of the focal point volume. The volume of dead cells is referred to as a lesion, and this method of selectively inducing cellular necrosis is commonly referred to as thermal ablation.
Some form of medical imaging can be employed to facilitate HIFU therapy. Because HIFU therapy can be carried out non-invasively, a non-invasive imaging technology will normally be employed. Medical imaging can facilitate HIFU therapy by enabling the exact location of abnormal tissue to be identified and targeted, by providing verification that the HIFU focal point properly corresponds to the abnormal tissue, by facilitating monitoring during HIFU treatment to ensure optimal therapeutic dose control, and by enabling post-therapy imaging of the treatment site to assess treatment efficacy. Ultrasound imaging for guiding HIFU therapy offers the following advantages: ultrasound imaging can be performed in real-time, ultrasound imaging has minimal side effects, ultrasound imaging equipment is ubiquitous and relatively inexpensive, ultrasound imaging and therapeutic ultrasound can be easily integrated into a single assembly, and acoustic distortions arising from sound wave and tissue interactions will generally have similar effects on both imaging ultrasound and HIFU.
Temperature measurements during HIFU therapy are useful for several reasons. First, a spatio-temporal temperature map overlaid over a B-mode ultrasound image of the treatment site will enable a clinician to prevent the temperatures of non-target tissue (such as tissue beyond the bounds of a tumor) from reaching necrotic levels. Further, temperature measurements can be used to calculate a thermal dose. It has been suggested and experimentally verified that, once a tissue type dependent thermal dose has been achieved for a given volume inside the tumor, irreversible damage will have been caused in this region. Based on the times for which temperatures are held, an equivalent time at one reference temperature (usually taken to be 43° C.) is computed, and this time estimate is referred to as the thermal dose. The mathematical expressions for thermal dose are as follows:
                                                        TD              43                        =                                          (                                  x                  ,                                                            t                      end                                        -                                          t                      0                                                                      )                            =                                                ∑                                      t                    =                                          t                      o                                                                            t                    end                                                  ⁢                                                      2                                                                  T                        ⁡                                                  (                                                      t                            ,                            x                                                    )                                                                    -                      43                                                        ⁢                  Δ                  ⁢                                                                          ⁢                  t                                                              ,                      T            >=            43                          ⁢                                  ⁢                                                            TD                43                            ⁡                              (                                  x                  ,                                                            t                      end                                        -                                          t                      0                                                                      )                                      =                                          ∑                                  t                  =                                      t                    o                                                                    t                  end                                            ⁢                                                0.5                                                            T                      ⁡                                              (                                                  t                          ,                          x                                                )                                                              -                    43                                                  ⁢                Δ                ⁢                                                                  ⁢                t                                              ,                      T            <            43                                              (        1        )            where T is the temperature at time t, Δt represents the time interval between consecutive temperature measurements, TD43 is the thermal dose referenced to 43° C., x is the spatial location in tissue where the thermal dose is computed, and to and tend represent the start time and end time of treatment, respectively. This expression has typically been used in hyperthermia treatments, in which the heating profile is relatively long and stays near 43.0° C., but also has been suggested for use in HIFU. From Eq. (1), it can be seen that estimating the temporal and spatial profile of the temperature rise during treatment is important for computing the thermal dose delivered to the patient.
Thus, it would be desirable to provide techniques and apparatus enabling ultrasound imaging and ultrasound data-based temperature estimates to be performed during HIFU therapy. Indeed, several suggested techniques have been investigated. In particular, the relationship between the speed of sound in tissue and temperature (i.e., acoustic attenuation) has been studied, in order to enable temperature monitoring to be achieved. Due to the nature of ultrasound imaging, it is relatively straightforward to collect time-of-flight data for ultrasound waves propagating between an imaging transducer and the treatment site. Changes in temperature at the treatment site will affect the speed at which reflected ultrasound waves propagate between the treatment site and the imaging transducer (acting as a receiver), enabling an estimation of temperature at the treatment site to be made.
Unfortunately, temperature estimates generated using the above-noted technique are prone to significant errors, for several reasons. Because the acoustic attenuation of a given mass of tissue is difficult to measure non-invasively, a generic acoustic attenuation value for biological tissue is generally employed to generate temperature estimates based on backscattered ultrasound data. However, the acoustic attenuation of biological tissue is not a constant value, and different types of tissue exhibit different acoustic attenuation (i.e., the speed of sound in fatty tissue is different than the speed of sound in muscle tissue, even at the same temperature). Thus, the use of a generic value for acoustic attenuation will introduce errors into the temperature estimates, and the use of numerical model-based approaches to derive temperature values assuming a priori knowledge of attenuation coefficient are therefore ineffective. It would therefore be desirable to provide a more accurate method and apparatus for obtaining profiles of the spatio-temporal temperature distribution of tissue during thermal therapy.